Field of the Invention
The present invention concerns a method that includes defining a diffusion-weighted magnetic resonance image, as well as a corresponding magnetic resonance apparatus. In particular, various embodiments relate to the variation in the amplitude and duration of diffusion-encoding gradient pulses.
Description of the Prior Art
Magnetic resonance (MR) imaging is a modality used for generating MR images based on MR data that depicts an examination subject. Typically, the examination subject, such as a person under examination, is positioned in a scanner in which a basic magnetic field is generated that is as static and as homogeneous as possible, with a field strength between 0.5 Tesla and 5 Tesla, for example. The basic magnetic field aligns the magnetization of nuclear spins the examination subject along the direction of the basic magnetic field.
Radio frequency (RF) excitation pulses are radiated into an examination region of the subject in order to deflect the nuclear magnetization from its neutral position along the direction of the basic magnetic field, that is, in order to excite nuclear magnetization. The subsequent relaxation of nuclear magnetization causes the nuclear spins to emit RF signals, known as echoes. In the context of gradient echo MR imaging or echo planar MR imaging (EPI), gradient echoes are generated, by using gradient pulses to rephase and dephase the nuclear magnetization (refocusing and dephasing gradient pulses). These can be part of a corresponding gradient pulse train, for example.
Gradient pulses can also be used for spatially encoding the MR data. The gradient pulses generate gradient magnetic fields (gradient fields), which are superimposed on the basic magnetic field.
The MR data are measured (detected) during a readout phase. The readout phase is chronologically spaced by the echo time TE against the excitation, so that the MR data are readout during one or more gradient echoes.
The acquired MR data are also referred to as raw data. The MR raw data can be processed in order to reconstruct the MR image of the examination subject. For example, the measured MR raw data are typically digitized and are initially available in the spatial frequency space (domain) (k-space). Using a Fourier transform, it is then possible to transform the MR raw data into image data in the image space, in order to generate the MR image.
A special form of MR imaging is diffusion-weighted MR imaging. In clinical routine, diffusion-weighted MR images can supply important diagnostic data, for example, in the diagnosis of strokes and tumors. Diffusion-weighted MR images contain information about the diffusion of molecules in the examination region. From the diffusion-weighting, it is possible to derive a degree of diffusion, so a quantitative statement can ensue.
In diffusion-weighted MR imaging, additional gradient fields (diffusion gradient fields) are produced in specific directions by activating corresponding diffusion-encoding gradient pulses. The diffusion gradient fields trigger the diffusion encoding of the MR data by diffusion: the diffusion of water molecules along the diffusion gradient fields typically attenuates the MR signal. In areas with a lower (higher) diffusion, a lower (higher) signal attenuation consequently typically ensues, such that, in an imaging MR measurement, these areas can have an increased (reduced) amplitude.
The degree of the diffusion weighting can be correlated with the strength of the diffusion gradient fields that are applied or with the amplitude and duration of the gradient pulses; typically stronger (weaker) diffusion gradient fields induce a higher (lower) diffusion weighting of the MR images.
The parameters of the diffusion gradient fields are often referred to a b value. b values for different gradient fields are known in relation to diffusion gradient fields and the so-called b matrix. The b matrix can describe properties of the diffusion gradient fields such as strength and/or orientation and/or duration etc.
The b matrix is determined on the basis of various b values. The b matrix is used to determine the diffusion tensor, which is a description of the degree and direction of diffusion. This can be achieved, for example, using an equation known as the Stejskal-Tanner equation. The diffusion tensor contains comprehensive information on the diffusion.
Different diffusion encodings are also possible. An example of diffusion encoding is dual bipolar diffusion encoding, cf. HEID O., “Eddy current-nulled diffusion weighting” in Proc. 8th Annual Meeting of ISMRM (2000) 799, the relevant disclosure from which is adopted here as a cross reference.
Known diffusion-weighted MR imaging techniques have certain limitations. In diffusion-weighted MR imaging, phase errors, such as slice-specific dephasing for example, may occur. This dephasing is caused by unwanted gradient fields that can occur as a consequence of the Maxwell equations. These unwanted gradient fields are also known as concomitant field terms.
Techniques for reducing phase errors due to concomitant field terms are known. For example, slice-specific and gradient-specific correction factors can be determined. These correction factors can be imprinted on the diffusion encoding by gradient pulses that have been adapted accordingly, cf. MEIER C. et al., “Concomitant Field Terms for Asymmetric Gradient Coils: Consequences for Diffusion, Flow, and Echo-Planar Imaging” in Mag. Reson. Med. 60 (2008) 128-134.
Such techniques as per MEIER C. can be complicated. In particular, it may be necessary to calculate the correction factors and adapt the gradient pulses accordingly.
Furthermore, such techniques may be useable for multi-slice MR imaging only with certain limitations. Multi-slice MR imaging is sometimes also referred to as slice multiplexed MR imaging. In multi-slice MR imaging, the nuclear magnetization is excited in two or more slices simultaneously and the MR data are readout simultaneously (i.e. simultaneous multi-slice, SMS). Separation of the MR data into the various slices is achieved, for example, on the basis of slice-specific reconstruction using a parallel image technique (partial parallel acquisition, PPA). PPA techniques make it possible to undersample k-space and to reconstruct missing sampling points in the MR data by using a reconstruction kernel. Corresponding techniques in relation to SMS EPI are described in: SETSOMPOP K. et al., “Improving diffusion MRI using simultaneous multi-slice echo planar imaging” in NeuroImage 63 (2012) 569-580, and in U.S. Pat. No. 8,405,395; these are incorporated as a cross reference here.
Since the aforementioned correction factors are slice-specific, it is not possible, or is only possible to a limited extent, to use them in the simultaneous modification of nuclear magnetization in multiple slices. Furthermore, in the context of SMS MR imaging, efforts are typically made to select the simultaneously modified slices as broadly as possible in order to allow a good slice separation or to avoid a significant lowering of the g factor. In the context of SMS MR imaging, spaced slices are selected in order to allow the separation of the slices by varying the sensitivities of the receiving coils. The farther apart the two simultaneously excited slices are, the greater is the variation in sensitivity. The g factor describes the lowering of the signal-to-noise ratio that is caused by the coil geometry. This is occurs particularly with high acceleration factors, that is, for a large number of simultaneously modified slices, or for RF coil arrays with a low number of coil elements. Therefore, it is often the case that a good approximation of the correction factors for all the simultaneously modified slices cannot be found. This may lead to a lowering of the quality of the MR images, in particular for slices that are situated at a distance from the isocenter of the scanner.
DE 10 2012 205 587 B4 discloses techniques in which RF excitation pulses for the slices are time-delayed. As a result, it is also possible in the context of SMS MR imaging to impose slice-specific factors that allow a correction of the concomitant field terms. However, such a technique has the limitation that different RF excitation pulses have to be determined for each slice, which may be complicated. Moreover, this may lead to a significant time-lag between the excitation of the nuclear magnetization in different slices, which may reduce the quality of the MR image.